The field of the invention is nuclear magnetic resonance imaging methods and systems. More particularly, the invention relates to a method and system for magnetic resonance imaging magnet assemblies.
Magnetic resonance imaging (MRI) is a diagnostic imaging modality that does not rely on ionizing radiation. Instead, it uses strong (ideally) static magnetic fields, radio-frequency (RF) pulses of energy and magnetic field gradient waveforms. More specifically, MR imaging is a non-invasive procedure that uses nuclear magnetization and radio waves for producing internal pictures of a subject. Three-dimensional diagnostic image data is acquired for respective xe2x80x9cslicesxe2x80x9d of an area of the subject under investigation. These slices of data typically provide structural detail having a resolution of one (1) millimeter or better.
When utilizing nuclear magnetic resonance (NMR) to produce images, a technique is employed to obtain NMR signals from specific locations in the subject. Typically, the region that is to be imaged (region of interest) is scanned by a sequence of NMR measurement cycles, which vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well-known reconstruction techniques. To perform such a scan, it is, of course, necessary to elicit NMR signals from specific locations in the subject. This is accomplished by employing gradient magnetic fields which have the same direction as a polarizing field, B0, but which are configured as needed to select the slice, phase encode and readout to facilitate the imaging. By controlling the strength of these gradients during each NMR cycle, the spatial distribution of spin excitation can be controlled and the location of the resulting NMR signals can be identified. When a substance such as human tissue is subjected to the uniform magnetic polarizing field, B0, the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in an x-y plane and which is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or xe2x80x9ctippedxe2x80x9d, into the x-y plane to produce a net transverse magnetic moment Mt. A signal is emitted by the excited spins after the excitation signal B1 is terminated, this signal may be received and processed to form an image.
When utilizing these signals to produce images, magnetic field gradients (GxGy and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
Unlike conventional magnets, an actively shielded magnet is unable to automatically compensate (i.e., via Lens""s Law) for the magnetic disturbances to the B0 field in the imaging volume due to external magnetic sources. This so because of the actively shielded magnet""s combination of positive and negative turns. Thus, the actively shielded magnet only partly compensates for the shift in the B0 field. A B0 coil is a secondary coil added to an actively shielded superconducting magnet to shield the effects of moving metal objects in the vicinity of the magnet. Unfortunately, B0 coils typically have a small mutual inductance with the primary coil and steadily accumulate current as the primary coil decays. This causes a deterioration in the magnetic field homogeneity in the imaging volume, leading to poor imaging quality. A quench in the primary coil can couple extremely high currents into the B0 coil, causing a risk of damage to the B0 coil.
The above discussed and other drawbacks and deficiencies are overcome or alleviated by a magnet assembly for a magnetic resonance imaging system. The magnet assembly comprises a primary coil including a first set of turns having a first prescribed number of turns about an axis. The first set of turns is symmetrically positioned radially from the axis and with respect to a mid plane perpendicular to the axis. A second set of turns has a second prescribed number of turns about the axis and is symmetrically positioned radially from the axis and with respect to the mid plane outward of the first set of turns.
A secondary coil includes a third set of turns having a third prescribed number of turns about the axis. The third set of turns is symmetrically positioned radially from the axis and with respect to the mid plane in close proximity to the first set of turns and outward of the first set of turns. A fourth set of turns has a fourth prescribed number of turns about the axis, and is symmetrically positioned radially from the axis and with respect to the mid plane in close proximity to the second set of turns and outward of the second and third sets of turns. The first and third sets of turns are in a first prescribed turns ratio and the second and fourth sets of turns are in a second prescribed turns ratio.
A method of optimizing the mutual inductance between a primary and a secondary coil in a magnetic resonance imaging actively shielded magnet assembly comprises minimizing deterioration in the homogeneity of the polarizing magnetic field for a given rate of change in the decay of the current in the primary coil; and minimizing the rate of change of current in the secondary coil and a shim coil for a given drift in the current in the primary coil.
The secondary coil shields the effects of moving metal objects in the vicinity of the magnet. Using the correct geometry, the B0 shielding function can be performed by a single secondary circuit which is non-coupling with respect to the primary coil. The primary and secondary circuits react to an external disturbance independently, according to Lens"" law, to completely cancel a shift in the B0 field in the imaging volume. The primary and secondary coils react independently of one another because there is no mutual inductance between them. Since the primary and secondary coils are non-coupling, the B0 coil does not accumulate current as a result of changes in the primary coil current. The B0 coil is divided into two parts which are electrically wired together in a series configuration. An inner coil is wound onto the positive turns of the primary coil and an outer coil is wound onto the negative turns (bucking coil) of the primary coil.
The above discussed and other features and advantages of the present invention will be appreciated and understood by those skilled in the art from the following detailed description and drawings.